Dendrimers are three-dimensional polymers that are grown by the successive addition of shells or layers of branched molecules to a central core. Dendrimers have several advantages over linear polymers, since they have controllable structure, a single molecular weight rather than a distribution of molecular weights, and a large number of controllable surface functionalities, and an inclination to adopt a globular conformation once a certain size is reached. They are prepared by reacting highly branched monomers together to produce monodisperse, tree-like and/or generational structure polymeric structures. Individual dendrimers consist of a central core molecule, with a dendritic wedge attached to each functional site. The dendrimeric surface layer can have a variety of functional groups disposed thereon, according to the assembly monomers used during the preparation. Generally, the dendrimer functional groups dictate the properties of the individual dendrimer types.
As a result of their design, dendrimer cores are spacious, and by modifying the chemical properties of the core, shells, and especially the surface layer, their physical properties can be finely tuned. Tunable properties include solubility, toxicity, immunogenicity and bioattachment capability. The core molecule is referred to as Generation 0. Each successive repeat unit along all branches, forms the next generation, Generation 1, Generation 2 and so on, until the terminating generation results. Half generations are possible, since generally, two condensation reactions steps are required to produce each full generation. Half generations are attained where generation formation is terminated at the first condensation reaction. Preparation of dendrimers requires a high level of synthetic control, which is achieved through series of stepwise reactions, which comprise building the dendrimer up one layer or generation at a time. Dendrimer synthesis can be of the convergent or divergent type. During divergent dendrimer synthesis, the molecule is assembled from the core to the periphery in a stepwise process involving attaching one generation to the previous and then changing functional groups for the next stage of reaction. Functional group transformation is necessary to prevent uncontrolled polymerisation. Such polymerisation would lead to a highly branched molecule that is not monodisperse and is otherwise known as a hyperbranched polymer. Hyperbranched polymers are undesirable, since they are thought to be potentially immunogenic. Due to steric effects, continuing to react dendrimer repeat units leads to a sphere shaped or globular molecule, until steric overcrowding prevents complete reaction at a specific generation and destroys the molecule's monodispersity. For this reason, typically dendrimers of G1-G10 generation are the most useful. However, the number of possible generations can be increased by reducing the spacing units in the branching polymer. For example, if ethyleneglycol is used during synthesis, generations greater than G10 can be prepared.
Dendrimers have two major chemical environments: namely specific surface chemistries, due to the functional groups on the termination generation (which is the surface of the dendritic sphere) and due to the sphere's interior (which is largely shielded from exterior environments due to the spherical shape of the dendrimer structure). The existence of two distinct chemical environments in such a molecule implies many possibilities for dendrimer applications. Dendrimers have found actual and potential use as molecular weight and size standards, gene transfection agents, as hosts for the transport of biologically important guests, in drug or biomolecule delivery and/or encapsulation systems, as micelles and as anti-cancer agents, to name but a few. Dendrimers' globular shape and molecular topology, however, make them highly useful to biological systems as well. In addition, dendrimers possess a strictly controlled number of functional groups on their periphery that can be used for the attachment.
Desirable properties for dendrimer use in biological systems include water solubility, lack of both toxicity and immunogenicity, low polydispersity, presence of internal voids and the presence of multiple, highly accessible functionalised arms for drug group, solubilising groups, targeting groups, or other moieties' attachment; all of which allow tuning of the biological properties of the system. Despite the large number of polymers available, relatively few possess all of these features. Many known dendrimers lack biodegradability, especially at higher molecular weights. This limits the size of dendrimer which can be used in biological applications.
Polyamidoamine (PAMAM), polypropyleneimine (PPI), polyarylether (PAE) and polyethyleneimine (PEI) are examples of dendrimers that have been investigated for biopharmaceutical applications [1-5]. Polyamidoamine (PAMAM) dendrimers are based on an ethylenediamine core and an amidoamine repeat branching structure. They can be synthesized in a variety of well-defined molecular weights. Their size and surface functionality (primary amine) is defined by the number of controlled repetitive additions of monomeric units, giving rise to different half or full generations. They are water-soluble and they have been reported to be the only class of dendrimer that are mono-dispersed. Furthermore, they show high charge densities that are restricted to the surface of the molecules. The synthesis of PAMAM can be tailored so as to influence the groups at the surface. Full PAMAM generations such as G1, G2, have amine functionalised surfaces to produce cationic dendrimers; while half Generations such as G1.5, G2.5, etc., have carboxylic acid groups at the surface and provide anionic dendrimers [6]. Interestingly, dendrimers adopt a tighter “native” shape or an extended “denatured” state, depending on solution pH.
Cationic PAMAM dendrimers, exhibit non-specific binding or interactions with cellular blood components and plasma proteins, such as albumin, fibronectin, immunoglobulins, complement factors and/or fibrinogen, have very short half-lives. Unfortunately, they have been shown to be toxic due to electrostatic interactions with the negatively charged components of cellular membranes [1]. They are known to cause destabilization of cell membranes and can result in cell lysis. In fact, in a recent study, PAMAM dendrimers have been shown not to be suitable carriers for agents such as biodegradable macromolecular MRI contrast agents, due to their high toxicity. Furthermore, regardless of the internal repeat unit structure, cationic dendrimers have been found to be haemolytic and cytotoxic, which is dependent on the generation and the number of surface ionic groups [2,7]. Efforts have been made to modify PAMAM dendrimer using polyethylene glycol to reduce these non-specific interactions and cytotoxicities. Hedden et al, grafted PEG chains of low polydispersity onto the terminal groups of PAMAM dendrimers [14]. Recently, Gillies et al, have reported a biocompatible polyester “bow-tie” dendrimer based on the monomer 2,2-bis(hydroxymethyl) propionic acid [12]. The new polyester dendrimer-PEG bow-tie hybrids were evaluated for their potential as drug delivery vehicles. In vitro experiments indicated that the polymers were nontoxic to MDA-MB-231 cells, and that they were degraded to lower MWs at the normal physiological pH of 7.4, and at the mildly acidic pH of 5.0. Other polyester based dendrimers are known, for example, U.S. Pat. No. 5,834,118 describes hyperbranched polyesters of a polyol with 3 to 10 reactive hydroxyl groups and an aromatic polycarboxylic anhydride with 2 to 4 carboxyl groups; each hydroxyl group of the polyol forms an ester linkage with one anhydride of the polycarboxylic anhydride, and further glycidyl (meth)acrylate or allyl glycidyl ether forming ester linkages with the remaining carboxylic group of the anhydride and free hydroxyl groups.
In 2003, International Publication No. WO 03/064502 described polyester dendrimers of polyols, wherein the dendrimer has 2 to 64 reactive hydroxyl groups. The polyester dendrimers are produced using alternating sequences of haloacetyl halide and carboxylate alkali metal salts. The dendrimer products are used as: cross linkers for polymers and gels, in glues and coatings, as matrix material for composites; as well as having useful roles in analytic and catalytic chemistry.
In 2004, Namazi et al. described biocompatible G1-G3 citric acid-polyethylene glycol-citric acid (CPEGC) triblock dendrimers, which were applied as solubilising controlled release drug-delivery systems for hydrophobic drugs, such as 5-amino salicylic acid (5-ASA), pyridine, mefenamic acid, and diclofenac [13]. This dendrimer has a polyethylene glycol core and N=1-3 generations comprising citric acid branch extenders.
The human body is not always capable of repairing or replacing damaged tissue. As a result, tissue structures or scaffolds have been developed in order to facilitate the regeneration of damaged or diseased tissue. Tissue scaffolds function to provide the support upon which cells can attach, grow and differentiate at the site of injury or damage.
Tissue engineering attempts to create three-dimensional tissue structures on which cells and other biomolecules may be incorporated. These structures or scaffolds guide the organization, growth, and differentiation of cells, in the process of forming functional tissue by providing physico-chemical cues. To successfully incorporate a scaffold within the host body depends on efficient communication between cells, tissues, and the host system as a whole. The scaffolds and cells to be incorporated must, therefore, interact with adhesion and growth factor receptors or bind such factors; and the scaffold must eventually degrade.
Many disease conditions or injuries of the body require the repair or replacement of damaged tissues, but the body itself may not be able to replace or repair the tissue satisfactorily, or do so within an appropriate time scale. Thus many methods of disease or injury treatment involve methods of augmenting the body's natural repair mechanisms and often rely on the use of implantable biological scaffolds or prostheses. Ideally, an implantable prosthesis should be chemically inert, noncarcinogenic, capable of resisting mechanical stress, sterilisable, and resistant to the actions of tissue fluids, as well as being non-inflammatory and hypoallergenic. A number of biological and synthetic scaffolds are known.
Scaffold substrates support growth and function of cells and, therefore, are essential to successful tissue engineering. Various tissues and tissue components of animal origin are currently used as scaffolds. Such scaffolds are desirably derived from natural polymer materials. Biological scaffolds when based on natural polymers, typically comprise proteins such as elastin, laminin, gelatin, fibrin, fibrinogen and collagen. Such biological scaffolds have a number of advantages over synthetic scaffolds: communication with existing body cells is instantaneous, they undergo a natural process of degradation, and existing biological signals attenuate incorporated signals (biological cues present in the scaffold provide signalling cues to the host cells) such as the growth factors and cytokines (which are inherently present in these scaffolds). Biological scaffolds must be noncarcinogenic, non-immunogenic, and biodegradable. Biological scaffolds must be cross-linked for strength and rigidity before implantation and use. However, such crosslinking modifications are not as easily made to biological scaffolds as they are to synthetic scaffolds. Minimally altered scaffolds will retain most of the natural molecules required to establish the expected therapeutic properties of the tissue-engineered organ or tissue. Upon implantation, these retained natural molecules are also further amenable to natural degradation.
Intact decellularised xenogenic/allogenic extracellular matrices (ECM), which are the parts of animal tissue that usually provide structural support to cells, are widely used in the manufacture of tissue-engineering scaffolds. The extracellular matrix (ECM) is essential for processes like growth, wound healing, and fibrosis. ECM comprises a matrix of natural polymers such as elastin, fibrin, laminin, fibronectin and collagen.
Advantages of utilising such matrices or individual natural polymers in biological scaffold applications arise from their low toxicity to tissue and their well-documented structural, physical, chemical, and immunological properties. In their purified state however, these polymers are weakly cross-linked and therefore do not possess the adequate mechanical properties needed for tissue engineering applications. The formation of covalent intermolecular cross-links between polymer molecules offers an effective method of improving the mechanical integrity and stability of natural polymers.
Polymer substrates have been widely used in biomaterials applications, for example: as wound dressings, matrices for the controlled release of active agents or as tissue engineering scaffolds, and more recently, in nanoshell applications such as targeted delivery of genes, antibodies, peptides and pharmaceuticals. Various procedures have been explored to cross-link natural polymers physically. Such methods include dehydrothermal treatment and ultraviolet irradiation to induce crosslinking between complementary groups within the polymer. Unfortunately, physical cross-linking, while avoiding the introduction of potentially cytotoxic chemical residues, leads to alteration of the chemical structure of the protein-based scaffolds (e.g. denaturation).
Use of polyethylene glycol (PEG) is of interest in production of biocompatible matrices. It is an interesting polyether for a number of reasons; PEG is a non-toxic, non-immunogenic, biodegradable molecule, which is water-soluble but yet will dissolve in organic solvents such as benzene and dichloromethane. PEG molecules can be linked to hydrophilic and/or hydrophobic drugs, allowing roles in drug delivery. However, to date, most cross-linking mechanisms using PEG are poorly controlled and not very efficient. Viers et al., attempted to overcome such deficiencies by endlinking telechelic vinyl sulfone (VS) end functionalised PEG with the terminal amines found on commercially available dendrimers such as PAMAM dendrimers. The VS terminated PEG endlinked to the dendrimer crosslinking molecules were successfully used to swell hydrogels.
A wide variety of other studies have reported the chemical use of the fixing agent glutaraldehyde (GA) as a means of cross-linking natural polymers. However, GA fixation is reportedly plagued by calcification and cytotoxicity problems. As a result, a significant amount of research has been directed towards finding alternative effective cross-linking mechanisms. Alternative chemical cross-linkers investigated include use of reagents such as diisocyanates, acyl azide and 1-ethyl-3-(3-dimethyl aminopropyl) carbodiimide (EDC). Most cross linking agents are “bridge-forming” agents, meaning that the polymer molecules are linked by a cross-linking molecule and the cross-links act as bridges between natural polymer molecules. Variation of the length of crosslinking side chains allows for control of spacing within the cross-linked polymer. EDC is also regularly used in organic chemistry to couple carboxylic acid groups to alcohol groups using, for example, (4-dimethylaminopyridine) DMAP as a catalyst. In biological applications, it is generally used as a carboxyl group activating agent for the coupling of primary amines to yield amide bonds. Additionally, it can also be used to activate phosphate groups towards derivatization. Thus, EDC has found roles in: peptide synthesis, protein crosslinking to nucleic acids, and preparation of immunoconjugates. EDC is often used in combination with N-hydroxysuccinimide (NHS) or sulfo-NHS to increase coupling efficiency or create a stable amine-reactive product.
During polymer EDC cross-linking, “zero-length” cross-links (wherein the natural polymer protein chains are directly bonded together) are formed, allowing the polymers to be directly linked to one another. An advantage of using the EDC method over other chemical cross-linking methods arises, because EDC leaves no cytotoxic residuals, and urea, the non-toxic by-product of this reaction, is easily washed off the treated natural polymer. A disadvantage, however, is that EDC crosslinking, using diamine from the EDC reagent as a cross linking bridge, reportedly offers little improvement in the biological stability of the resultant material.
Of particular interest is the recent use of dendrimers in natural polymer crossing linking. Their potential use as a candidate for crosslinking stems from the ability of the surface functional groups to react with natural polymer groups, such as amine, amide, hydroxyl or thio functionalities, to form covalent bonds and to act as a bridging linker that provides for good spacing within the natural polymer fibres.
Recently several publications have disclosed the use of branched polymer molecules, such as dendrimers, as a means of cross-linking natural polymers. Sheardown et al., describe a method of cross-linking collagen polymers with an amine functionalised polyamidoamine (PAMAM) dendrimer using carbodiimide (EDC) chemistry. This dendrimer has an amine surface functional group available for cross-linking with the carboxyl group of amino acids of natural polymers. However, for cross-linking to take place using the PAMAM/EDC system, the natural polymer functional groups have to be first activated using an in situ carbodiimide reaction. This results in the activation of the carboxylic (COOH) groups of the collagen polymer towards the amine (NH2) groups of the PAMAM dendrimer. However, this system suffers from severe shortcomings; the collagen carboxylic group activation mechanism can lead to internal cross-linking between the internal COOH groups and NH2 groups of collagen natural polymer to result in the production of zero-length cross links. This is undesirable since internal collagen crossing linking of this type has the effect of reducing the space available for cell penetration into the polymer matrix or scaffold, and also, may restrict the passage for oxygen and nutrients thereto.
Despite the drawback of zero-length cross links, the use of functionalised dendrimers as natural polymer cross-linkers is preferable over use of other cross linking agents; advantageously, they can be utilised to provide a variety of chemical functionalities to natural polymeric systems. The availability of functional groups on the surface of the dendrimer permits the binding of, for example, exogenous bioactive molecules such as drugs, markers or endogenous molecules such as hormones or antibodies. Thus, the provision of biological functionalities within the scaffold, that act to guide tissue regeneration, is highly desirable. Currently available technologies using carbodiimides (such as those described above) are further limited in that they will not permit the fuctionalisation of the polymer matrix without altering the biological properties of the polymer itself. Furthermore, it will be appreciated that the stability of a biological scaffold or polymer matrix can be tailored by controlling the extent of cross-linking. This is easily achieved through adjusting the natural polymer, dendrimer cross-linking ratios. Therefore the amount of dendrimer used in the cross-linking reaction is critical in controlling scaffold degradation; a lesser degree of cross-linking will mean the polymer is easier to degrade.
There is still a need to provide new dendrimer systems that are highly biocompatible, are biodegradable, have well-defined MWs, and have functional groups. More desirably still, there is a need to provide an improved dendrimer having all of the aforementioned characteristics, but in addition, one that is more reactive than existing functionalised dendrimers; ideally the improvement will negate the requirement for toxic chemicals, such as EDC, for use in dendrimer applications.